Ultrasound is widely used in medical diagnosis for imaging of tissues, e.g. human fetus in utero, and measuring the velocity of blood and heart motion by means of Doppler shift of the backscattered ultrasound. In such applications, it is desirable that the ultrasound transducer is as sensitive as possible, thus maximizing the signal energy used to produce the image or measure the tissue velocity. As the available transducer materials that generate and receive ultrasound signals are of different acoustic impedances than human tissue, “matching” between the tissue and the transducer is necessary to maximize the signal that is received. For this reason, use of a “matching layer” is known to those skilled in the art of making ultrasound transducers. Using a transducer matching layer is the application to ultrasound devices of the same theory as employed in optics for anti-reflection coatings or for matching networks in electronics: increasing the amount of energy that enters a desired medium from driving source by reducing the energy reflected back caused by a mismatch in impedances between the driving impedance and the propagating medium's impedance.
It can be shown (for example, in G. Kino, Acoustic Waves: Devices, Imaging and Analog Signal Processing, Prentice-Hall, 1987, page 12) that to maximize the amount of acoustic energy of wavelength λ from a transducer of impedance Zt into a medium of impedance Zm, a layer is required with the characteristics of being one-quarter wavelength thick and having an acoustic impedance of (Zt·Zm)1/2, i.e. of impedance equal to the square root of the product of the transducer impedance and the medium impedance. This is the same relation as needed to match electromagnetic transmission lines.
A matching layer is widely used in ultrasonic transducers for medical imaging, for example, as taught by Utsumi et al in U.S. Pat. No. 4,756,808. This is because of the large impedance difference between the most commonly used piezoelectric material for medical ultrasonic transducers such as the ceramic PZT (Lead Zirconate Titanate), and human tissue. The unit of acoustic impedance is Rayl—named after Lord Rayleigh—in units of kg/m2-sec. PZT has an impedance of 20-40×106 rayl (i.e. 20-40 MRayl), while the acoustic impedance of tissue is 1.5 MRayl, leading to poor transfer of energy into tissue, as needed for medical imaging, unless a matching layer is used. Therefore, most medical imaging transducers have a matching layer attached to their outer surface to increase the transducer's sensitivity by improving coupling between PZT and human tissue.
Piezoelectric materials include a number of piezoplastic materials, e.g. PVDF, Nylon 8, vinylcyanide-vinylacetate copolymer, of which the most acoustically efficient is P(VDFx−TrFE100-x), a copolymer. When x, the percentage of VDF in the copolymer, is in the range of 65-82 mol % which is the desirable range for transducer operation (See Ohigashi, et al, Piezoelectric and Ferroeectric properties of P[VDF−TrFE] Copolymers and Their Application to Ultrasonic Transducers, in Medical Applications of Piezoelectric Polymers, Ed. by P. Galletti et al, Gordon and Breach Science Publishers, New York, 1988), the copolymer can be dissolved and spin-coated or dipped onto a substrate, and after annealing and polarizing, will form a light, flexible, piezoelectric film. While not as efficient a transducer material as ceramic PZT (kT2, the measure of conversion of electrical to acoustic energy for copolymer is about ⅓rd that of PZT), the copolymer has the advantages of ease of fabrication, lightness, and flexibility that make it desirable in certain situations. Moreover, because of the copolymer's low planar coupling, individual transducers elements can be defined by placement of metal electrodes, in contrast to PZT, whose high planar coupling requires mechanical grooves to be cut between elements to allow them to function independently. Particularly, the ease of making transducer elements by, for example, simple photolithographic deposition of metal electrodes rather than by mechanical grooving of micron-size cuts is an important advantage at higher frequencies, e.g. from 5 MHz to 30 MHz, where the dimension of the individual transducer elements needed to direct an ultrasound beam must be a fraction of the 50-500 micron acoustic wavelength. Furthermore, it is hard to fabricate these very thin, e.g. 250-25 micron thick for the 5-50 MHz range, ceramic transducers.
Another advantage of the P(VDF−TrFE) copolymer material is that its acoustic impedance is about 4.5 MRayl, much closer to the impedance of water (or tissue) of 1.5 MRayl than PZT's ˜35 MRayl. Note that while the mismatch is much smaller, there is still a substantial 3:1—mismatch between the impedances, so a matching layer would still improve the coupling between such a copolymer piezoplastic transducer and tissue.
However, while this copolymer has been used as a transducer for two decades, matching layers have been rarely used. At the high frequencies for which piezoplastic transducers are particularly advantageous, attaching such layers to a piezoplastic layers with minimal bond thickness—so as not to affect the matching—is extremely difficult. Therefore, although there has been development of piezoplastic transducers, such as taught in U.S. Pat. No. 6,641,540 to Fleisschman et al, or U.S. Pat. No. 8,156,620 to Habu et al, matching layers are not included in their fabrication.
With reference to FIG. 1A, in the conventional structure of a piezoplastic ultrasonic transducer 100, a layer of piezopolymer 130, is spin-coated (or dipped, or electro sprayed etc, as is known in the art) onto a substrate, 110, on which there is an electrode layer 120. A second electrode layer, 121, is formed on top of the piezopolymer to complete the transducer.
Recently M. Toda taught in published US patent applications 2002/0027400 and 2011/0050039 ways of combining polymer films and metal films to provide composite matching layers for piezoceramic or piezoplastic resonant transducers. This provides means of synthesizing composite layers of desired acoustic characteristics. Toda teaches using a polymer layer attached to the piezo-element upon which is placed a metal layer between the polymer-element of the synthesized matching layer and the medium into which the acoustic energy is to propagate. With reference to FIG. 1B, a matching layer 140 was added by means of an adhesive layer 135, for example as discussed by Toda. However, the addition of a thin bonding layer reduces the flexibility of the structure and increases the difficulty of fabrication and cost of the transducer. An improved structure and fabrication method of transducer is further needed.